This invention relates generally to magnetic resonance imaging (MRI), and more particularly, the invention relates to characterizing and correcting spatial gradient non-uniformities in diffusion tensor imaging.
Magnetic resonance imaging (MRI) requires placing an object to be imaged in a static magnetic field, exciting nuclear spins in the object within the magnetic field, and then detecting signals emitted by the excited spins as they precess within the magnetic field. Through the use of magnetic gradient and phase encoding of the excited magnetization, detected signals can be spatially localized in three dimensions.
FIG. 6A is a perspective view partially in section illustrating conventional coil apparatus in an NMR imaging system, and FIGS. 6B-6D illustrate field gradients which can be produced in the apparatus of FIG. 6A. This apparatus is discussed by Hinshaw and Lent “An Introduction to NMR Imaging: From the Bloch Equation to the Imaging Equation” Proceedings of the IEEE, Vol. 71, No. 3, March 1983, pp. 338-350. Briefly, the uniform static field B0 is generated by the magnet comprising the coil pair 10. A gradient field G(r), with r=[x y z]T, is generated by a complex gradient coil set which can be wound on the cylinder 12. An RF field B1 is generated by a saddle coil 14. A patient undergoing imaging would be positioned within the saddle coil 14.
In FIG. 6B an X gradient field is shown which is parallel to the static field B0 and ideally varies linearly with distance along the X axis but ideally does not vary with distance along the Y or Z axes. FIGS. 6C and 6D are similar representation of the Y gradient and Z gradient fields, respectively.
FIG. 7 is a functional block diagram of conventional imaging apparatus as disclosed in NMR-A Perspective in Imaging, General Electric company. A computer 20 is programmed to control the operation of the NMR apparatus and process FID signals detected therefrom. The gradient field is energized by a gradient amplifier 22, and the RF coils 26 for impressing an RF magnetic moment at the Larmor frequency are controlled by the transmitter 24. After the selected nuclei have been flipped, the RF coil 26 is employed to detect the FID signal which is passed to the receiver 28 and thence through digitizer 30 for processing by computer 20.
The use of diffusion tensor magnetic resonance imaging (DTI) for imaging anistropic tissue such as brain white matter is known. See U.S. Pat. No. 6,463,315 for analysis of Cerebral White Matter for Prognosis and Diagnosis of Neurological Disorders as well as U.S. Pat. No. 5,539,310 for Method and system for measuring the diffusion tensor and for diffusion tensor imaging and art cited therein. DTI provides a novel way to characterize tissues based on sensitivity to microscopic molecular motion of water. Clinical implementation requires strong, fast hardware and careful post processing of diffusion parameters. Diffusion weighted images and derivatives such as the three principal diffusivities of the diffusion tensor are quite specific in reflecting the physical properties of diffusion.
Diffusion weighted imaging (DWI), in general, consists of estimating the effective scalar diffusivity of water, D, in each voxel from a set of diffusion weighted images. During the time of a typical magnetic resonance data acquisition, water molecules diffuse on the order of a few microns, which is comparable to the dimensions of cellular structures, but significantly less than the dimensions of a voxel. Since D is sensitive to the physical properties, composition and spatial distribution of the tissue constituents, the measurement is sensitive to the tissue microstructure and physiological state.
Diffusion along a given axis is typically measured by placing a pair of diffusion sensitizing gradient pulses in the same axis in the magnetic resonance (MR) pulse sequence. The gradient pulses impose position-dependent phases on water protons that are equal in magnitude but opposite in sign and therefore cancel for stationary spins. However, for protons that move between the two gradient pulses, a finite net phase is accumulated. The sum of all phases from all protons results in attenuation of the MR signal due to interference effects. The magnitude of signal attenuation is dependent on the diffusivity of water, and the width, separation and amplitude of the gradient pulses. In a generalized case where the diffusivity may differ in different directions, a diffusion tensor matrix notation is used.
Non-uniformities of magnetic field gradients can cause serious artifacts in diffusion tensor imaging. While it is well appreciated that non-linearities of the imaging gradients lead to image warping (see for example Bernstein et al., U.S. Pat. No. 6,163,152), those imperfections can also cause spatially dependent errors in the direction and magnitude of the diffusion encoding.